1. Technical Field
Embodiments of the invention relate generally to system and method for medical imaging. Particular embodiments relate to systems and methods for attenuation correction of combined positron emission tomography/magnetic resonance imaging (“PET-MR”).
2. Discussion of Art
Positron emission tomography (“PET”) machines use one or more rings of scintillators or other detectors to generate electrical signals from gamma rays (photon pairs) that produced from the recombination of electrons, within a target material, and positrons, emitted from decay of a radionuclide packaged in a tracer compound. Typically, recombination events occur within about 1 mm from the radionuclide decay event, and the recombination photons are emitted in generally opposite directions to arrive at different detectors. Paired photon arrivals that occur within a detection window (usually less than a few nanoseconds apart) are counted as indicating a recombination event, and, on this basis, computed tomography algorithms are applied to the scintillator position and detection data in order to locate the various recombination events, thereby producing three-dimensional images of the tracer disposition within the target material.
Typically, the target material is body tissue, the tracer compound is a liquid analogue to a biologic fluid, and the radionuclide is disposed primarily in body tissues that make use of the biologic fluid. For example, a common form of PET makes use of fludeoxyglucose (18F), which is analogous to glucose with the 18F radionuclide substituted for one of the hydroxyl groups ordinarily composing glucose. Fludeoxyglucose is preferentially absorbed by brain matter, by the kidneys, and by growing cells, e.g., metastasizing cancer cells. As a result, PET is frequently used for oncologic studies, for localizing particular organs, and for studying metabolic processes.
One challenge in obtaining desired PET image quality is that gamma rays, in the energy spectrum produced by positron-electron interactions, are easily attenuated by typical body tissues and are differently attenuated by different body tissues. Varying attenuation can diminish statistical confidence in the locations of recombination events, thereby making the computed image “fuzzier” than is desirable. Accordingly, it is highly desirable to provide means for attenuation correction (“AC”).
For example, “phantoms” (artificial targets of known shape and material properties) can be imaged for evaluating PET performance, calibrating PET imaging software, and developing product capability specifications. Often, PET is combined with computed tomography (“CT”), which uses a moving X-ray source and detectors to obtain images of internal structures. X-rays are photons and are attenuated much like the lower-energy photons produced from positron-electron recombination events. Thus, AC algorithms, using CT data, can be useful for improving PET image quality in real time. Phantom scans in a PET-CT machine can be done using the same parameters and protocol as for a patient scan. The machine does not need adjustment to account for the phantom not being a patient.
Magnetic resonance imaging (“MR”) uses magnetic sensors to detect rotating magnetic fields that are produced by nuclei that have odd atomic numbers, i.e., total number of neutrons and protons is not divisible by two, in response to alternately imposing and removing a magnetic field. Typically, MR is accomplished by imposing on a target material a strong, e.g., 1.5-3 Tesla, magnetic field that pulses or fluctuates at radio frequencies, e.g., 900 MHz. Field strength is important for establishing a steep field gradient, which helps in determining the location of nuclei. Radio frequency is important for amplifying the rotating fields, which are produced each time the fluctuating field is imposed. MR can also be accomplished using low-frequency, low-strength fields such as the geomagnetic field.
MR is frequently used for differentiating tissue types within a patient, and is also used for identifying fine detail structures. Typically, different pulse sequences are used for tissue differentiation. For example, a T1 pulse sequence can be used to obtain images with water appearing darker and fat brighter. On the other hand, a T2 pulse sequence can be used to obtain an image with fat darker, and water lighter.
One advantage of MR is that magnetic fields do not attenuate in body tissues, so that nucleus location can be determined (using Fourier analysis) based solely on frequency shifting between the imposed magnetic field and the response field. Another advantage is that by careful selection of pulse sequence, distinct tissues or materials can be highlighted.
PET and MR can be combined in a single apparatus (a “PET-MR scanner”). Such an apparatus provides fine detail, tissue differentiation, and metabolic data. However, because MR signals do not attenuate in the same way as PET or CT signals attenuate, and because MR signal return is highly dependent on the type of pulse sequence used (with each pulse sequence emphasizing a different material, whereas the PET signal is attenuated by every material intervening between a recombination event and a pair of detectors), it becomes difficult to obtain reliable AC of the PET signals based on MR signals from a single scan. For example, an MR image that is obtained using pulse sequences chosen to highlight radionuclide inserts within a phantom fill fluid, typically will not return a signal for the solid material of a phantom casing, which attenuates the PET signal from the fill fluid. Essentially, the MR image data is not useful for correcting the attenuated PET image data. Moreover, a schedule of MR pulse sequences, optimized for clinical scans (patient body materials) may not reliably return a usable image from a standard PET phantom.
Some of the difficulties in AC of PET images, in a PET-MR machine, can be overcome in case exact position of a phantom is known. Then prior knowledge of the phantom materials and structure can provide an ideal image, an AC parameters can be tuned to approximate the ideal image from the actually-obtained phantom image. This approach is similar to a manual registration technique in which a calculated attenuation map is manually registered to acquired data and used for reconstruction. Manual positioning of phantoms, or registration of an attenuation map with acquired data, however, is extremely difficult to accomplish with the precision needed for accurate AC.
Additionally, and more significantly, PET signal attenuation is very specific to patient anatomy and material properties. As already discussed, an MR image that does display the positron-absorbing phantom fluid, often does not display the photon-absorbing phantom casing. This problem is generalizable to clinical practice: Standard phantoms do not have the identically non-uniform distribution of tissue as is found in a patient. As such, attenuation of a phantom scan is not expected to accurately map to a subsequent patient scan, and phantom-based AC is not implemented using the same protocol as would be used in clinical settings for a patient scan.
As a result, other techniques for attenuation correction have been proposed. Many of these techniques depend upon receiving a locationally accurate MR signal from the liquid filling a PET phantom. For example, one approach involves mixing PET tracer in a solvent optimized for MR. The fluid body within a PET phantom can, however, exhibit MR image distortion for various reasons. Additionally, some phantoms are filled with a semi-solid gel, which may not return a quality MR image. Other problems with implementing this approach relate to the material specificity of MR pulse sequences.
In view of the above, aspects and embodiments of the present invention provide attenuation correction of a PET phantom image, in a PET-MR machine, without relying on the MR data.